Images of the interiors of bodies may be acquired using various types of tomographic techniques, which involve recording and measuring radiation from tissues and processing acquired data into images.
One of these tomographic techniques is positron emission tomography (PET), which involves determining spatial distribution of a selected substance throughout the body and facilitates detection of changes in the concentration of that substance over time, thus allowing to determine the metabolic rates in tissue cells.
The selected substance is a radiopharmaceutical administered to the examined object (e.g. a patient) before the PET scan. The radiopharmaceutical, also referred to as an isotopic tracer, is a chemical substance having at least one atom replaced by a radioactive isotope, e.g. 11C, 15O, 13N, 18F, selected so that it undergoes radioactive decay including the emission of a positron (antielectron). The positron is emitted from the atom nucleus and penetrates into the object's tissue, where it is annihilated in reaction with an electron present within the object's body.
The phenomenon of positron and electron annihilation, constituting the principle of PET imaging, consists in converting the masses of both particles into energy emitted as annihilation photons, each having the energy of 511 keV. A single annihilation event usually leads to formation of two photons that diverge in opposite directions at the angle of 180° in accordance with the law of conservation of the momentum within the electron-positron pair's rest frame, with the straight line of photon emission being referred to as the line of response (LOR). The stream of photons generated in the above process is referred to as gamma radiation and each photon is referred to as gamma quantum to highlight the nuclear origin of this radiation. The gamma quanta are capable of penetrating matter, including tissues of living organisms, facilitating their detection at certain distance from object's body. The process of annihilation of the positron-electron pair usually occurs at a distance of several millimeters from the place of the radioactive decay of the isotopic tracer. This distance constitutes a natural limitation of the spatial resolution of PET images to a few milimeters.
A PET scanner comprises detection devices used to detect gamma radiation as well as electronic hardware and software allowing to determine the position of the positron-electron pair annihilation event on the basis of the position and time of detection of a particular pair of the gamma quanta. The radiation detectors are usually arranged in layers forming a ring around object's body and are mainly made of an inorganic scintillation material. A gamma quantum enters the scintillator, which absorbs its energy to re-emit it in the form of light (a stream of photons). The mechanism of gamma quantum energy absorption within the scintillator may be of dual nature, occurring either by means of the Compton's effect or by means of the photoelectric phenomenon, with only the photoelectric phenomenon being taken into account in calculations carried out by current PET scanners. Thus, it is assumed that the number of photons generated in the scintillator material is proportional to the energy of gamma quanta deposited within the scintillator.
When two annihilation gamma quanta are detected by a pair of detectors at a time interval not larger than several nanoseconds, i.e. in coincidence, the position of annihilation position along the line of response may be determined, i.e. along the line connecting the detector centers or the positions within the scintillator strips where the energy of the gamma quanta was deposited. The coordinates of annihilation place are obtained from the difference in times of arrival of two gamma quanta to the detectors located at both ends of the LOR. In the prior art literature, this technique is referred to as the time of flight (TOF) technique and the PET scanners utilizing time measurements are referred to as TOF-PET scanners. This technique requires that the scintillator has a time resolution of a few hundred picoseconds.
Currently, the state of the art methods of determining the sites of interactions of the gamma quanta in positron emission tomography are based on the measurements of charges of signals generated in vacuum tube photomultipliers, silicon photomultipliers, or avalanche diodes optically connected to inorganic crystals notched into smaller elements. Position of the gamma quantum reaction is determined with the accuracy of the size of a smaller crystal element on the basis of the differences in charges of the signals from different converters optically connected to the same crystal. In the state of the art PET scanners, reconstruction of the set of LOR and TOF data is based on the relationships between charges and times of signals recorded for a particular event without reference to external reference signals.
In the signal time determination methods used in the state of the art, changes in shapes and amplitudes of signals depending on the place of ionization and the quantity of energy constitute a limitation in temporal resolutions that can be achieved using the technique. The larger the scintillator, the larger the variations in signal shapes and amplitudes.
For the above reasons, temporal resolutions of less than 100 ps are unattainable in the state of the art for large scintillator blocks. Temporal resolution also translates on the resolution of ionization place determination. In case of polymer scintillators (preferred due to their low price), amplitudes of signals generated by the gamma quanta, including annihilation gamma quanta used in positron emission tomography, are characterized by continuous distribution resulting from interactions between gamma quanta and electrons occurring mostly via the Compton effect with a negligibly low probability of a photoelectric effect. As a consequence, signal amplitudes in polymer scintillators may change even if they originate from the same place.
As shown by the shortcomings of the state of the art signal analysis techniques, there is a need to significantly improve temporal and spatial resolution in the detectors used in medical diagnostic techniques that require recording of ionizing radiation. The need to improve resolution is particularly high in large-sized detectors. Examples of PET detectors making use of large polymeric scintillators were described in the PCT application WO 2011/008119 as well as in the PCT application WO 2011/008118. Solutions described in these applications are based on the measurements of the times of light pulses arrival to the detector edges. Light pulses are converted into electric pulses by means of photomultipliers. The shape (distribution of photons as a function of time) and the amplitude of the light pulse reaching the photomultiplier varies depending on the distance between the photomultiplier and the pulse origin place. In addition and independently of the ionization place, the amplitude of the signal varies with the energy deposited within the detector. As a consequence, due to variations in signal shapes and amplitudes, it is impossible to achieve good temporal resolution using either leading edge or constant fraction discriminators used in state of the art, due to the time walk effect and the pulse shape change effect observed in large-size scintillators.
The goal of this invention is to develop a method for reconstructing the place of the reaction of gamma quanta in PET detectors as well as for reconstructing the difference between the times of flight (TOF) of the annihilation quanta to different detectors such that said method would not deteriorate spatial or temporal resolution capabilities even in cases when the times of arrival and the shapes of the recorded pulses would vary greatly depending on the place of reaction of a particular gamma quantum within the detector.